Conventional procedures based on two-dimensional gel electrophoresis for profiling the concentrations of specific proteins and their byproducts are time-consuming, labor-intensive, and require significant technical expertise to obtain quantitative information. One approach for circumventing these limitations is to develop the equivalent of a DNA microarray for proteins. Protein microarrays consist of various types of capture ligands that exhibit a high binding affinity toward a particular protein. Target proteins are either labeled with a fluorescent label, or additional fluorescently labeled protein is used to selectively bind to the target once it has been capture to a specific site on the array. The latter approach, often referred to as a sandwich assay, has the advantage of being extremely selective since very rarely will a target protein bind to both the capture molecule and an additional protein. This disadvantage of the sandwich assay is that there are a limited number of target proteins for which there exist two distinct binding partners. As a result, direct labeling of the target protein can be most generally applied for profiling multiple proteins from a cell lysate. However, there are two drawbacks of direct labeling: First, the efficiency of coupling fluorescent labels to low abundance proteins within a lysate is highly variable. This can often make it difficult to achieve sufficient sensitivity as well as reliability. Second, the complexity of protein structures poses significant challenges for attaching labels to specific sites while preserving the functionality of a protein. This challenge was cogently summarized by biophysicist S. P. Fordor: “Conventional detection techniques based on fluorescent tagging require that one partner of a complex is chemically modified. These modifications can subtly alter molecular interactions by changing the chemical nature of the binding interaction.” Mazzoila, L. T. and Fordor, S. P. A., Biophys. J. (1995), 68:1653-1660, the entire teachings of which are incorporated herein by reference. Thus, eliminating the labeling process will improve the feasibility, speed, and utility of quantitative protein assays.
The major limitation of existing label-free detectors is that they are significantly less sensitive than fluorescence detection. The two most well known approaches are the quartz crystal microbalance (QCM) for detecting surface adsorbed mass and the surface plasmon resonance (SPR) technique for detecting refractive index changes in close proximity to a metal surface. Both methods have significant fundamental limitations concerning scalability, sensitivity to low-concentration samples, and their ability to provide quantitative information. The mass resolution of the QCM is on the order of 10−17 g/μm2, which corresponds to about 100 proteins/μm2 (assuming a molecular weight 100 kDa). Furthermore, the QCM sensor area is macroscopic in scale (typically a few mm2), so the minimum detectable mass is on the order of several nanograms, or 1010 molecules. This detection level is not suitable for many biological assays. Fluorescence routinely resolves 1-10 molecules for a surface area less than 100 μm2. The QCM also requires that the capture ligands be rigidly coupled to the sensor surface. This limits the efficiency of three-dimensional coatings (e.g. carboxymethyldextran (CMD) matrix) that enhance the effectiveness of mass sensing.
The SPR, which achieves a similar resolution as the QCM, measures changes in the refractive index that occur within a CMD layer above the sensor surface several hundred nanometers thick. Since the influence of the target molecules on the optical properties of this layer is generally unknown, SPR usually provides indirect information. This can make it difficult to quantify the amount of bound target molecules. Furthermore, attempts to reduce the sensor surface area for large-scale integration are not yet capable of reaching sensitivity levels that are comparable to commercial macroscopic instruments. This limitation is often attributed to the difficulty of matching the very narrow operating range of the integrated optics to the refractive index of typical buffer solutions using materials available for microfabrication.
The development of label-free detectors that are both sensitive and scalable (both down in sensor area and up in number of sensors) is in its infancy. In addition to on-going research for advancing optical methods as well as acoustical methods such as the flexural plate wave device (FPW), there are several new approaches for label-free detection that are currently being pursued. One approach for molecular detection is the transduction of surface binding events on a microcantilever into mechanical bending. The bending is not induced by the addition of mass but rather the change in surface energy resulting from specific binding of the biomolecules. For example, it has been shown that a microcantilever stress sensor can detect DNA hybridization. In other work, it has been shown that the microcantilever stress sensor can detect prostate-specific antigen (PSA) in a background of human serum albumin and human plasminogen (Wu et al., Nature Biotech. (2001), 19:856, the entire teachings of which are incorporated herein by reference). This result suggests that the stress sensor could be a clinically relevant diagnostic technique for prostate cancer.
The advantage of using microcantilevers for label-free detection is that this technique is scalable. For example, researchers at IBM Zurich have demonstrated that eight cantilevers can be detected in parallel (Arntz, et al., Nanotechnology (2003), 14(1): 86 and Battiston, et al., Sensors and Actuators B-Chemical (2001), 77(1 and 2): 122-131, the entire teachings of which are incorporated herein by reference). However, there are three drawbacks to using the microcantilever stress sensor. First, there is not yet a viable approach for integrating the cantilevers with conventional microfluidics. Currently, the surfaces must be functionalized by manually aligning micropipettes to each cantilever before packaging the cantilevers within a macroscopic fluid cell. Second, the surface stress induced by molecular binding must occur on only one side of the cantilever in order for it to bend. This requires different surface chemistries to be developed for the top and bottom sides of the sensor. Third, it has not yet been demonstrated (or predicted) that the stress sensor has a higher resolution (in terms of minimum number of detectable molecules per area) than commercially available sensors such as the SPR. Despite these limitations, the development of the microcantilever stress sensor is still in its infancy and its full potential has yet to be realized.
Finally, resonant cantilever mass sensors, while successful for chemical sensing in gaseous environments, have received less attention for biomolecular detection in solution. This is primarily because the mass sensitivity and frequency resolution are significantly degraded by the low quality factor and large effective mass that is induced by viscous drag. While the quality factor can be enhanced by using electronic feedback known as Q-control, the mass sensitivity, in terms of frequency shift per mass loading, is not improved.
Accordingly, there remains a need for an analytical technique that is sufficiently sensitive but does not require modification of the analyte of interest.